In the past, a radiation-sensitive scintillation layer is employed for converting a ray into a light signal, a television camera is employed for receiving the light signal, and a display screen is employed for displaying the light signal, so that real-time imaging may be realized during radioscopy. With the development of the technologies, the emergence of CCD makes such a real-time imaging mode more optimized. CCD has very superior performances on stability, integrity, consistency and high-speed collection. However, due to the problem of radiation damage resistance of CCD itself, such a real-time imaging mode has an inevitable defect, that is, a radiation sensitive film with a sufficient thickness or a device for the light transmitting path is needed. The commonly used conversion/transmission devices comprise an image enhancer, a lens and an optical fiber, etc., and these conversion/transmission devices are located in front of the CCD in the work flow, thus the factors such as inconsistency, noise and so on that are introduced thereby make various advantages of CCD unable to be shown sufficiently, and at the same time, the complexity of the imaging system is increased and the reliability of the system is lowered.
Since 1990s, technical personnel skilled in the field of radiation imaging has started a study on how to combine advantages such as high speed, high image quality, high integrity, high reliability, large area and operation simplicity, etc., on a radiation imaging detector, thereby developing a digital image flat-panel detector with a large area.
At present, a large area of thousands of square centimeters, a spatial resolution of tens of micrometers and a reading speed of hundreds of frames per second may be realized by a product of digital image flat-panel detector.
A first type of prior art of digital image flat-panel detector employs a combination of an amorphous silicon diode and a TFT. The amorphous silicon diode absorbs a ray and generates electron-hole pairs. Under the influence of an electric field, charged particles with a certain polarity drift to a TFT pixel array, and each pixel signal is read sequentially by the switched scanning with respect to the TFTs.
The ionization energy of amorphous silicon is very low (about 5 eV), accordingly, a great number of electron-hole pairs can be generated under the irradiation of a ray, and a good signal-to-noise ratio can be obtained even in a low dosage.
However, the atomic number of silicon is very small (Z=14) and its ability to barrier the ray is rather weak, thus a very thick silicon layer is needed to effectively barrier the ray. This makes the method difficult to be realized technically and makes the cost very high.
A second type of prior art of digital image flat-panel detector employs a combination of an amorphous selenium film and a TFT. The amorphous selenium absorbs a ray and generates electron-hole pairs. Under the influence of an electric field, charged particles with a certain polarity drift to a TFT pixel array, and each pixel signal is read sequentially by the switched scanning with respect to the TFTs.
The atomic number of selenium is 34 and its ability to barrier a ray is stronger than that of the amorphous silicon, but it is only applicable for detecting a ray below 50 KeV. This limits the main application area of an amorphous selenium flat-panel detector to a low energy field (for example, Mammography).
The ionization energy of the amorphous selenium changes as the strength of applied field and the energy of the incident ray vary. In the ranges of the field strength and the ray energy commonly used in medical diagnosis, the ionization energy thereof is about 50 eV, thereby the lowest dosage and the output signal amplitude of the ray are limited.
Moreover, the temperature stability of the amorphous selenium is poor, and it is easy to be deliquesced and crystallized. Thus, its life time is not as good as flat-panel detectors with other structures.
A third type of the prior art of digital image flat-panel detector employs a combination of a scintillator, a photodiode and a TFT. The scintillator converts a ray into a light signal, the photodiode receives the light signal and converts it into an electric signal, and then each pixel signal is read sequentially by the switched scanning with respect to the TFTs.
The scintillator may absorb the ray energy and emit photons of visible light with a wavelength in a certain range, and the number of the photons emitted is in proportion to the energy absorbed. The atomic number of the scintillator material is generally high, and its ability to absorb the ray is strong. The scintillator may be a fluorescent film material (for example, certain rare earth materials) or a scintillating crystal (for example, cesium iodide, cadmium tungstate, etc.).
The atomic number of cesium iodide crystal is larger than either amorphous silicon or amorphous selenium, and it has a good barrier and absorbing ability on rays. Meanwhile, the emission spectrum peak position of thallium-doped cesium iodide crystal is 565 nm, which basically fits the absorption spectrum peak position of the amorphous silicon photodiode, and a combination of the cesium iodide crystal and the amorphous silicon photodiode has the highest quantum efficiency among the products of the same type. Due to these advantages, most of the current digital image flat-panel detectors have a structure in which a cesium iodide crystal, a silicon photodiode and a TFT are combined.
When the scintillator is a homogeneous film material, in order to increase the detectable energy range and the detection efficiency, it is needed to increase the thickness of the film. However, as the film thickness increases, the influence of the scattering of optical photons on the spatial resolution of the detector increases. When the scintillator is a cesium iodide crystal, the scattering of photons may be suppressed by growing the crystal so as to form a high-density acicular array (a needle tube with a size of 10-20 μm).
However, as the thickness of the cesium iodide film increases, the aspect ratio of the needle tube increases, and the collection efficiency of photons inside the tube lowers greatly, so that the quantum efficiency of the detector is decreased. At the same time, due to the problem of size matching between the needle tube and the photodiode, the proportion occupied by the dead zone of such a detector is prone to be large.
Inside the scintillator, the generation of each optical photon requires energy of about 20-50 eV; moreover, in consideration of the quantum efficiency of the photodiode on the visible light wave band, a detector with such a structure requires the energy of about 100 eV or even more to generate each electron-hole pair. Such a performance determines that a relatively poor signal-to-noise ratio will be obtained when a scintillator is employed as a radiation sensitive film.
A fourth type of prior art of digital image flat-panel detector employs a combination of a scintillator and a CMOS. The scintillator may be directly overlaid on the CMOS, or an optical fiber with different diameters on its two ends may be employed to combine a scintillator having a large area with a CMOS having a small area.
By substituting a CMOS process for the traditional silicon process, the system integrity may be increased to a greater extent, and the spatial resolution, duty ratio, collection speed and so on of the detector may be increased greatly. Each pixel unit is integrated with an independent charge-voltage converting circuit and an independent amplifying circuit, thus a better signal-to-noise ratio can be obtained.
However, as limited by the CMOS process, it is difficult for such a flat-panel detector to obtain a large sensitive area under a low cost. But, it has evident advantages in the small area detection field, for example, dentistry CT and CT for small animals, etc.
In conclusion, in the prior art digital image flat-panel detectors, when factors such as dynamic range (detectable energy range), detection efficiency, signal-to-noise ratio and spatial resolution, etc., are considered, the main ray conversion mode is as follows: a ray is first converted into an photon of visible light by using a high-density acicular cesium iodide scintillator, and then the photon of visible light is converted into an electric signal via a photodiode.
TFT reading or CMOS reading is mainly employed for the reading of an electric signal. One of the TFT and CMOS is selected according to the actually required factors such as the area, spatial resolution, collection speed, integrity, cost and so on.